Organic electrolyte gated field-effect transistor biosensor

Abstract

The organic electrolyte-gate field-effect transistor biosensor includes a microfluidic channel structure formed from a dielectric thermoplastic material. A biorecognition body immobilized within the microfluidic channel enables the biosensor to detect the presence or concentration of an analyte in the electrolyte fluid placed within the channel. The dielectric material separates the microfluidic channel from the gate electrode and from the semiconductor material connecting the drain and source electrodes, preventing direct contact between the gate electrode and the semiconductor material and the electrolyte fluid within the microfluidic channel, thus providing the biosensor with high capacitance. A method for fabricating the biosensor by monolithic 3D printing is also provided.

Description

[Technical Field]

[0001] field This application relates to biosensors. More specifically, this application relates to improved organic electrolyte gate field-effect transistor biosensors, as well as related methods and uses for measuring the presence and quantity of analytes in biological and other samples. [Background technology]

[0002] background According to the definition of the International Union of Pure and Applied Chemistry (IUPAC), a biosensor can be considered a specific type of chemical sensor. A chemical sensor is thought to consist of two basic functional units: a receptor and a transducer. When an analyte interacts with the receptor, the transducer converts chemical information about the analyte, such as its presence and concentration, into an analytically useful signal. In a biosensor, the interaction of the receptor with the analyte occurs through biochemical mechanisms such as binding to or reactions with proteins, polynucleotides, or other biological molecules (e.g., antibodies, antigens, enzymes, receptors, aptamers, etc.), or through interactions between the analyte and biological substances such as cells or tissues.

[0003] One type of transducer used in chemical or biosensors is the field-effect transistor (FET). FETs are semiconductor devices and traditionally include inorganic semiconductor materials such as silicon, often doped with small amounts of elements such as boron, aluminum, gallium, indium, phosphorus, arsenic, or antimony, but organic semiconductor materials can also be used. Commonly used organic semiconductor materials are characterized by aromatic, conjugated, or delocalized π-electron systems and include compounds such as rubrene, pentacene, and thiophene polymers, which may be further chemically substituted to adjust their conductive properties. Semiconductor materials can be classified as n-type, where electrons are the primary charge carriers, or p-type, where holes are the primary charge carriers.

[0004] A FET has three terminals or electrodes, designated as gate, source, and drain. A conductive channel in the semiconductor material connecting the source and drain allows current to flow between the source and drain terminals. However, applying a voltage to the gate can change the conductivity of the channel and, consequently, the output current of the FET. If the gate voltage changes in relation to the interaction between the receptor of a chemical or biosensor and the analyte, the resulting change in the output current of the FET can provide an analytically useful signal that reflects information related to the analyte.

[0005] For example, in insulated-gate FETs such as MOSFETs (metal-oxide-semiconductor FETs), the presence of an insulating dielectric layer between the gate and the conductive channel allows the FET to behave as a capacitor. Depending on the voltage applied to the gate, an electric field is generated, which can increase or decrease the number of charge carriers in the conductive channel depending on the specific composition of the FET, thereby increasing or decreasing the output current flowing between the source and drain terminals. Therefore, for example, if the gate is positively polarized with respect to the conductive channel, the semiconductor material of the conductive channel becomes electron-rich, the gate electrode becomes positively charged net, while the semiconductor becomes negatively charged net. If the conductive channel contains an n-type semiconductor, the number of electrons available to carry charge increases. Similarly, if the gate is negatively polarized with respect to the conductive channel, the semiconductor material of the conductive channel becomes electron-depleted, the gate electrode becomes negatively charged net, while the semiconductor becomes positively charged net. If the conductive channel contains a p-type semiconductor, the electron depletion will result in an increase in the number of holes available to carry charge. In either of these cases, increasing the gate voltage will result in an increase in the number of charge carriers in the conduction channel, and therefore an increase in the output current flowing between the source and drain terminals. Conversely, decreasing the gate voltage will result in a decrease in the number of charge carriers in the conduction channel, and therefore a decrease in the output current flowing between the source and drain terminals.

[0006] Electrolyte-gate FETs, or EGFETs, in their basic form, include an electrolyte as part of the dielectric layer between the gate electrode and the semiconductor material. The electrolyte may be, for example, an aqueous ion-containing solution, an ionic liquid, an ionic gel, or a conductive polymer. When a bias voltage is applied to the gate, ions in the electrolyte move toward or away from the gate, depending on their charge. Therefore, for example, if the gate is negatively polarized with respect to the semiconductor material as described above, positively charged cations from the electrolyte are attracted to the negatively charged gate and move toward it, resulting in the formation of an electric double layer at the interface between the electrolyte and the gate. Similarly, negatively charged anions in the electrolyte are repelled by the negatively charged gate and move toward the positively charged semiconductor, resulting in the formation of a second electric double layer at the interface between the electrolyte and the semiconductor surface. Alternatively, if the gate is positively polarized with respect to the semiconductor material as described above, negatively charged anions from the electrolyte will move toward the positively charged gate, and positively charged cations in the electrolyte will move toward the negatively charged semiconductor, again resulting in the formation of an electric double layer at each interface.

[0007] The presence of an electric double layer enables the formation of a high degree of charge separation between the gate and the conductive channel for a given applied gate voltage, thus providing the EGFET with a high gate capacitance. It is noteworthy that the capacitance of the EGFET is proportional to the ion concentration. Therefore, a change in the ion concentration of the gate electrolyte can result in a corresponding change in the gate capacitance in the resulting electric field at a given gate voltage, and in the conductivity of the conductive channel. Thus, such a change in ion concentration can be reflected in a measurable change in the output current of the EGFET.

[0008] EGFETs, including organic semiconductor materials (EGOFETs), are well-suited for use as transducers in biosensors because at least one can be easily incorporated into a small, flexible device that can be efficiently manufactured using 3D printing technology. As a result of high gate capacitance, EGOFETs can operate at lower applied gate voltages, reducing the likelihood of causing electrolysis of aqueous media and thus making them more compatible with biological systems. In addition, smaller changes in gate voltage or capacitance result in larger changes in conductivity and output current of the conductive channel, leading to increased sensitivity. Furthermore, EGOFETs can utilize biological fluids such as blood, plasma, urine, sweat, breath condensate, or saliva as the gate electrolyte and include a layer at the interface between the electrolyte and either the gate or the semiconductor that allows for selective binding of the analyte. It has been found that when the analyte is present in the biological fluid, binding to the functionalized layer can cause changes in gate capacitance, resulting in a change in output current that reflects the concentration of the analyte in the fluid. As a result, EGOFETs may be particularly useful in biosensors designed to detect the presence or concentration of a target analyte in aqueous biological fluids.

[0009] Previous reports attempting to design biosensors using EGOFETs have shown that while it is possible to prepare biomolecular-sensitive devices that demonstrate conversion changes reflecting changes in analyte concentration, the reported devices exhibit limited reproducibility and high variability in correlating the output from complex samples with analyte concentration. One possible complicating factor is the contribution of pseudocapacitance to total capacitance, resulting from charge transfer between the electrolyte and gate electrode due to electroadsorption or intercalation processes, or reversible Faraday redox reactions. Furthermore, known processes used to passivate the surface of semiconductor materials can allow fluid penetration from the electrolyte to the semiconductor surface, leading to degradation of the surface passivation layer, resulting in loss of sensitivity and reduced device shelf life.

[0010] Organic electrolyte gated FET (OEGFET) biosensors that overcome some of these shortcomings have been reported. For example, see below: Massey, R., et al, "Label-free detection of dopamine using aptamer enhanced organic-electrolyte gated FET sensor." In 2019 IEEE International Conference on Flexible and Printable Sensors and Systems (FLEPS), pp. 1-3. IEEE, 2019, Massey, R., et al, "Aptamer-Enhanced Organic Electrolyte-Gated FET Biosensor for High-Specificity Detection of Cortisol." IEEE Sensors Letters 4, no. 7 (2020): 1-4, Massey, R., et al, "A Comprehensive Modelling Approach for Bio-EDLC systems." In 2020 IEEE Sensors, pp. 1-4. IEEE, 2020, R. Massey et al, “A System-On-Board Integrated Multi-Analyte PoC Biosensor for Combined Analysis of Saliva and Exhaled Breath”, 2022 IEEE International Engineering in Medicine and Biology Society Conference, pp. 1-6, 2022, R. Massey et al, “System-On-Board Integrated Flexible OEGFET Aptasensor for Saliva Testing of Cortisol”, 2022 IEEE International Conference on Flexible and Printable Sensors and Systems (FLEPS), pp. 1-4, 2022、および、 R. Massey et al, “A Low Temperature Processed, Soft-fluidic OEGFET Saliva Aptasensor for Cortisol”, IEEE Journal on Flexible Electronics, 1, no. 1, pp. 64-72, Jan. 2022。

[0011] For example, as described in "RS Massey and R. Prakash, "A Low-Temperature-Processed, Soft-Fluidic OEGFET Saliva Aptasensor for Cortisol," in IEEE Journal on Flexible Electronics, vol. 1, no. 1, pp. 64-72, Jan. 2022," these OEGFETs incorporate a two-plane design in which the microfluidic channel structure is sandwiched between an upper and lower layer. The upper layer is prepared by using a shadow mask to deposit a chromium gate electrode onto a flexible polyamide film (Kapton®), followed by a layer of poly(methyl methacrylate) (PMMA). The microfluidic channel structure is molded or stamped from polydimethylsiloxane (PDMS) and cured and sealed into the PMMA upper layer using a crosslinking process involving (3-aminopropyl)triethoxysilane (APTES). Next, an aptamer designed to recognize a specific analyte is drop-casted onto the PMMA surface of the upper layer using UV activation, so that the liquid present in the microfluidic channel is exposed to the biosensing surface of the upper layer. The lower layer is prepared by depositing aluminum and chromium onto a Kapton™ film using a shadow mask to form the source and drain electrodes. An organic semiconductor such as 6,13-bis(triisopropylsilylethynyl)pentacene (TIPS-pentacene) is spin-welded to the surface and annealed to connect the source and drain electrodes, and covered with multiple stacked layers of dielectric material such as polyvinyl alcohol (PVA) and an outer layer of uncured PDMS that contacts the microfluidic channel and is sealed by curing. Thus, both the gate electrode and the semiconductor connecting the source and drain electrodes are shielded from contact with the electrolyte fluid in the microfluidic channel by the dielectric material, thereby reducing or eliminating the potential effects of pseudocapacitance and avoiding damage to the semiconductor from exposure to the electrolyte fluid.

[0012] OEGFET-based biosensors fabricated using this method have been reported to exhibit useful concentration-dependent conversion changes for analytes such as cortisol, dopamine, and α-synuclein. However, these biosensors exhibit several drawbacks. For example, the lack of control over the crystal growth direction of some organic semiconductors, which leads to variations in the level of resistance to electron transport through the conductive channel and inconsistencies in performance between different devices, prevents the fabrication process from scaling up to a commercial level and results in process variability. Furthermore, since the microfluidic channels are molded from PDMS, their thickness and filling can vary from device to device, further contributing to inconsistencies. The need to crosslink the PDMS microfluidic channels to PMMA dielectric material also leads to incomplete sealing of the microchannels and the possibility of subsequent leakage of the electrolyte sample. In addition, the shelf life of the devices is not sufficient for viable commercial use. Moreover, the fabrication process is not sufficiently suitable for incorporating biorecognition molecules other than aptamers into the biosensing surface. Therefore, a biosensor that mitigates or overcomes some of these drawbacks is desired. [Overview of the project]

[0013] overview In one aspect, the present application provides an organic electrolyte gate field-effect transistor biosensor comprising a gate electrode, a source electrode, a drain electrode, and a semiconductor material in electrical contact with the source electrode and the drain electrode. The biosensor also comprises a microfluidic channel structure formed from at least one thermoplastic dielectric material and comprising at least one inner surface comprising a biorecognition body, wherein the at least one inner surface is separated from the gate electrode, source electrode, drain electrode, and semiconductor material by the thermoplastic dielectric material. In at least one embodiment, the biorecognition body is an aptamer. In at least one embodiment, the biorecognition body is an antibody.

[0014] In at least one embodiment, the microfluidic channel structure includes a first dielectric layer having a first surface, a second dielectric layer having a second surface parallel to and opposite to the first surface, and one or more sidewalls adjacent to the first and second dielectric layers, thereby separating the first dielectric layer from the second dielectric layer and defining a microfluidic channel between the first surface, the second surface, and one or more sidewalls. The microfluidic channel is configured to receive an electrolyte fluid such that the electrolyte fluid is in contact with the first and second surfaces. In addition, at least one of the first and second surfaces constitutes a biorecognition body. In at least one embodiment, the first surface constitutes a biorecognition body. In at least one embodiment, the second surface constitutes a biorecognition body. In at least one embodiment, both the first and second surfaces constitute a biorecognition body.

[0015] In at least one embodiment, the semiconductor material is an organic or carbon-based semiconductor material. In at least one such embodiment, the semiconductor material is selected from the group consisting of 6,13-bis(triisopropylsilylethynyl)pentacene (TIPS-pentacene), polythiophenes and their copolymers, metal phthalocyanines, and single-walled carbon nanotubes. In at least one such embodiment, the semiconductor material comprises 6,13-bis(triisopropylsilylethynyl)pentacene (TIPS-pentacene) or carbon nanotubes. In at least one embodiment, the semiconductor material is a metal oxide semiconductor material. In at least one embodiment, the semiconductor material is formed by depositing it onto a self-assembled monolayer. In at least one embodiment, the self-assembled monolayer comprises octadecyltrichlorosilane.

[0016] In at least one embodiment, the thermoplastic dielectric material of the microfluidic channel structure, and in at least one embodiment, the thermoplastic dielectric material of the first and second dielectric layers and one or more sidewalls, is one or more materials selected from the group consisting of polylactic acid (PLA), polycaprolactone (PCL), poly(ethyl methacrylate) (PEMA), polybutylene terephthalate (PBT), polyethylene terephthalate (PET), polyethylene (PE), polypropylene (PP), polystyrene (PS), polytetrafluoroethylene (PTFE), cellulose, acylated cellulose derivatives, cellulose acetate, cellulose acetate butyrate, cellulose propionate, acrylonitrile butadiene styrene (ABS), acrylonitrile styrene acrylate (ASA), polycaprolactam (PC), polybutadiene (PBD), polydimethylsiloxane (PDMS), and nylon polyamide. In at least one embodiment, the thermoplastic dielectric material comprises at least one of polylactic acid (PLA) and polycaprolactone (PCL).

[0017] In at least one embodiment, one or both of the first dielectric layer and the second dielectric layer contain at least one of polylactic acid (PLA) and polycaprolactone (PCL). In at least one embodiment, the first dielectric layer and the second dielectric layer each contain polylactic acid (PLA), polycaprolactone (PCL), or a mixture of PLA and PCL. In at least one embodiment, both the first dielectric layer and the second dielectric layer each contain PCL. In at least one embodiment, both the first dielectric layer and the second dielectric layer each contain PLA. In at least one embodiment, the first dielectric layer contains PCL and the second dielectric layer contains PLA. In at least one such embodiment, the first surface constitutes the biorecognition body. In at least one embodiment, the first dielectric layer contains PLA and the second dielectric layer contains PCL. In at least one such embodiment, the second surface constitutes the biorecognition body.

[0018] In at least one embodiment where one or both of the first dielectric layer and the second dielectric layer contain PCL, the biometric body can be directly bonded to the PCL on one or both of the first surface and the second surface. In at least one embodiment where one or both of the first dielectric layer and the second dielectric layer contain PCL, the biometric body can be bonded to a layer of poly(methyl methacrylate) (PMMA) on one or both of the first surface and the second surface.

[0019] In at least one embodiment where the first dielectric layer and the second dielectric layer each contain PLA, the first dielectric layer contains a polylactic acid - polyethylene glycol - carboxylic acid block copolymer, and the first surface bears the biometric body. In at least one embodiment, the second dielectric layer contains a polylactic acid - polyethylene glycol - carboxylic acid block copolymer, and the second surface bears the biometric body.

[0020] In at least one embodiment, at least one of the first dielectric layer and the second dielectric layer includes a layer containing PLA, a layer containing polyvinyl alcohol (PVA) and nanocrystalline cellulose, and a layer containing PCL.

[0021] Another aspect of the present application provides a method for producing an organic electrolyte gate type field effect transistor biosensor as described herein. In at least one embodiment, the method includes the following: 3D printing a thermoplastic dielectric material to form a microfluidic channel structure including at least one inner surface; Placing the microfluidic channel structure between a gate electrode and a semiconductor material in electrical contact with a source electrode and a drain electrode such that at least one inner surface is separated from the gate electrode and the semiconductor material by the dielectric material; and, Bonding a biometric body to at least a part of at least one inner surface, including. In at least one embodiment, the microfluidic channel structure is three-dimensionally printed as a monolithic structure from a thermoplastic dielectric material.

[0022] In another aspect, this application provides a method for analyzing an electrolyte fluid for an analyte, comprising exposing an organic electrolyte gate field-effect transistor biosensor as described herein to the electrolyte fluid and determining one or more characteristics of the analyte. In at least one embodiment, the electrolyte fluid is a sample obtained from an environmental source. In at least one embodiment, the electrolyte fluid is a sample obtained from an agricultural source. In at least one embodiment, the electrolyte fluid is a sample obtained from a food source. In at least one embodiment, the electrolyte fluid is a sample obtained from a manufacturing source. In at least one embodiment, the electrolyte fluid is a biological fluid. In at least one embodiment, the biological fluid is one or more of blood, plasma, urine, sweat, interstitial fluid, cerebrospinal fluid, saliva, or exhaled condensate. In at least one embodiment, the electrolyte fluid is saliva. In at least one embodiment, the electrolyte fluid is blood. In at least one embodiment, the electrolyte fluid is plasma.

[0023] In at least one embodiment, the analyte is a biomarker whose presence or concentration indicates a medical condition. In at least one embodiment, the analyte in the electrolyte fluid is selected from the group consisting of drugs, metabolites, hormones, neurotransmitters, enzymes, carcinogens, peptides, proteins, electrolytes, metal ions, nucleic acids, and cells. In at least one embodiment, the analyte in the electrolyte fluid is an analyte useful for measuring nerve function, cardiac function, liver function, or kidney function, or for detecting neurodegenerative diseases, heart disease or cardiac disorders, blood diseases or blood disorders, infectious diseases, inflammation or stress-related diseases, liver disease or liver disorders, kidney disease or kidney disorders, and cancer or precancerous symptoms. In at least one embodiment, the analyte in the electrolyte fluid is selected from the group consisting of α-synuclein, β-amyloid, and tau protein.

[0024] In at least one embodiment, exposure of the biosensor to an electrolyte fluid includes exposure of the microfluidic channels of the biosensor to the electrolyte fluid. In at least one embodiment, determining one or more features of the analyte includes determining the presence of the analyte in the electrolyte fluid or measuring the concentration of the analyte in the electrolyte fluid. In at least one embodiment, the interaction of the analyte with the biorecognition body results in the formation of a charge separation region at the interface between the electrolyte fluid and at least one of the first dielectric layer and the second dielectric layer, a change in the net capacitance of the gate-to-channel of the biosensor, and a change in the output current of the biosensor, where the change in output current indicates one or more features of the analyte in the electrolyte fluid. [Brief explanation of the drawing]

[0025] Brief explanation of the drawing Further features of the present invention will become apparent from the following description and the accompanying drawings, which are not necessarily drawn to scale.

[0026] [Figure 1] Figure 1 is a schematic top view illustrating an embodiment of the biosensor described in this application.

[0027] [Figure 2] Figure 2 is a cross-sectional view along line 2-2 in the schematic representation of Figure 1.

[0028] [Figure 3A] Figure 3A is a schematic cross-sectional view illustrating an alternative embodiment of the biosensor described in this application.

[0029] [Figure 3B] Figure 3B is a schematic cross-sectional view illustrating an alternative embodiment of the biosensor described in this application.

[0030] [Figure 4A] Figure 4A is a schematic cross-sectional view illustrating an alternative embodiment of the biosensor described in this application.

[0031] [Figure 4B] Figure 4B is a schematic cross-sectional view illustrating an alternative embodiment of the biosensor described in this application.

[0032] [Figure 5] Figure 5 is a schematic cross-sectional view illustrating an alternative embodiment of the biosensor described in this application.

[0033] [Figure 6A] Figure 6A is a schematic front view illustrating an alternative embodiment of the biosensor described in this application.

[0034] [Figure 6B] Figure 6B is a cross-sectional view along line BB of the schematic representation in Figure 6A.

[0035] [Figure 6C] Figure 6C is a cross-sectional view along line CC in the schematic representation of Figure 6A.

[0036] [Figure 7A] Figure 7A is a schematic cross-sectional view illustrating an alternative embodiment of the biosensor described in this application.

[0037] [Figure 7B] Figure 7B is a schematic cross-sectional view illustrating an alternative embodiment of the biosensor described in this application.

[0038] [Figure 8A] Figure 8A is a schematic top view illustrating an alternative configuration of the microfluidic channel in the biosensor described in this application.

[0039] [Figure 8B] Figure 8B is a schematic top view illustrating an alternative configuration of the microfluidic channel in the biosensor described in this application.

[0040] [Figure 8C]Figure 8C is a schematic top view illustrating an alternative configuration of the microfluidic channel in the biosensor described in this application.

[0041] [Figure 8D] Figure 8D is a schematic top view illustrating an alternative configuration of the microfluidic channel in the biosensor described in this application.

[0042] [Figure 9A] Figure 9A schematically illustrates the binding modes in which the biorecognition body of the biosensor embodiment described in this application can bind to the analyte.

[0043] [Figure 9B] Figure 9B schematically illustrates alternative binding modes in which the biorecognition body of the biosensor embodiment described in this application can bind to the analyte.

[0044] [Figure 9C] Figure 9C schematically illustrates another alternative binding mode in which the biorecognition body of the biosensor embodiment described in this application can bind to the analyte.

[0045] [Figure 10A] Figure 10A is a graph showing the variation in device output current (IDS) with respect to input channel voltage (VDS) at various applied gate voltages (VGS) for this biosensor, in which the aptamer acts as the biorecognition component, when deionized water is present as the electrolyte.

[0046] [Figure 10B] Figure 10B is a graph showing the variation in the device output current (IDS) for applied gate bias (VGS) at various drain-source voltages (VDS) in the absence of an electrolyte, in order to measure the device transfer function in the embodiment of Figure 10A.

[0047] [Figure 10C]Figure 10C is a graph showing the variation in device output current with respect to input channel voltage (VSD) at various applied gate voltages (VGS) for a biosensor that contains polycaprolactone (PCL) as a dielectric material, carries an antibody as the biorecognition body, and is used in the presence of deionized water as an electrolyte.

[0048] [Figure 10D] Figure 10D is a plot showing the variation in channel current, transconductance (transfer function), and capacitance with respect to the gate modulation voltage (VSG) in the embodiment of Figure 10C.

[0049] [Figure 10E] Figure 10E is a graph showing the variation of the device output current (ISD) with respect to the input channel voltage (VSD) at various applied gate voltages (VGS) for a biosensor that contains polylactic acid (PLA) as a dielectric material, carries an antibody as the biorecognition body, and is used in the presence of deionized water as an electrolyte.

[0050] [Figure 10F] Figure 10F is a plot showing the variation in channel current, transconductance (transfer function), and capacitance with respect to the gate modulation voltage (VSG) in the embodiment of Figure 10E.

[0051] [Figure 11] Figure 11 is a graph showing the variation in device current with respect to various channel voltages (VDS) for this biosensor, in which semiconductor channels are prepared by various methods.

[0052] [Figure 12] Figure 12 is a graph showing the variation in device current with respect to channel voltage when exposed to synthetic buffers containing cortisol at various concentrations, for a biosensor in which a cortisol-recognizing aptamer serves as the biorecognition body.

[0053] [Figure 13] Figure 13 is a graph showing the variation in device current with respect to channel voltage when exposed to a synthetic buffer containing α-synuclein at various concentrations, for a biosensor in which an aptamer that recognizes α-synuclein (α-Syn) serves as the biorecognition body.

[0054] [Figure 14] Figure 14 is a plot showing the device current and capacitance of this biosensor, in which an aptamer that recognizes α-synuclein (α-Syn) serves as the biorecognition body, when exposed to saliva containing α-synuclein at various concentrations.

[0055] [Figure 15A] Figure 15A is a graph showing the variation in device output current (ISD) with respect to input channel voltage (VSD) at various applied gate voltages (VGS) for a biosensor configuration containing single-walled carbon nanotubes as semiconductors and polycaprolactone (PCL) as dielectric material, in the presence of deionized water as the electrolyte.

[0056] [Figure 15B] Figure 15B is a graph showing the variation in device output current (IDS) with respect to applied gate voltage (VSG) at various drain-source voltages (VSD) to measure the device transfer function of the embodiment shown in Figure 15A in the presence of deionized water as the electrolyte.

[0057] [Figure 15C] Figure 15C is a graph showing the variation of the device output current (ISD) with respect to the input channel voltage (VSD) at various applied gate voltages (VSG) in the embodiment of Figure 15A, in the presence of tris-acetic acid-EDTA (TAE) buffer as the electrolyte.

[0058] [Figure 15D]Figure 15D is a graph showing the variation in device output current (ISD) with respect to applied gate voltage (VSG) at various drain-source voltages (VSD) to measure the device transfer function of the embodiment shown in Figure 15A in the presence of TAE buffer as an electrolyte.

[0059] [Figure 16A] Figure 16A is a plot showing the device current and capacitance of a biosensor, which contains polycaprolactone (PCL) as a dielectric material and uses an aptamer that recognizes α-synuclein as the biorecognition body, when exposed to TAE buffer solutions containing α-synuclein at various concentrations.

[0060] [Figure 16B] Figure 16B is a plot showing the device current and capacitance of an alternative embodiment of this biosensor, which contains polycaprolactone (PCL) as a dielectric material and uses a monoclonal antibody that recognizes amyloid-β-42 as the biorecognition body, when exposed to TAE buffer solutions containing amyloid-β-42 at various concentrations.

[0061] [Figure 16C] Figure 16C is a plot showing the device current and capacitance in the configuration of Figure 16A when exposed to genuine salivary supernatant (RSN) samples containing various concentrations of α-synuclein.

[0062] [Figure 16D] Figure 16D is a plot showing the device current and capacitance in the configuration of Figure 16B when exposed to RSN samples containing various concentrations of amyloid-beta-42.

[0063] [Figure 16E]Figure 16E is a plot showing the device current and capacitance in the configuration of Figure 16A when exposed to real serum samples (referred to as "WTS") obtained from mice and containing various concentrations of α-synuclein.

[0064] [Figure 16F] Figure 16F is a plot showing the device current and capacitance in the configuration of Figure 16B when exposed to WTS serum samples containing various concentrations of amyloid-β-42. [Modes for carrying out the invention]

[0065] Detailed explanation In one aspect, this application provides an organic electrolyte gate field-effect transistor biosensor comprising a gate electrode, a source electrode, a drain electrode, and a semiconductor material in electrical contact with the source electrode and the drain electrode. The biosensor also includes a microfluidic channel structure formed from a dielectric material and comprising at least one inner surface that constitutes the biorecognition body.

[0066] In at least one embodiment, the microfluidic channel structure includes a first dielectric layer having a first surface, a second dielectric layer having a second surface parallel to and opposite to the first surface of the first dielectric layer, and one or more sidewalls adjacent to the first and second dielectric layers, separating the first surface from the second surface and defining a microfluidic channel between them. In at least one embodiment, the first dielectric layer separates the gate electrode from the microfluidic channel. In at least one embodiment, the second dielectric layer separates the semiconductor material from the microfluidic channel. In at least one embodiment, the first dielectric layer separates the gate electrode from the microfluidic channel, and the second dielectric layer separates the semiconductor material from the microfluidic channel, thereby separating the gate electrode from the semiconductor material by the first dielectric layer, the microfluidic channel, and the second dielectric layer. This arrangement can provide a high capacitance between the gate electrode and the semiconductor material in an organic electrolyte gate field-effect transistor biosensor.

[0067] Figures 1 and 2 schematically represent such an embodiment of the biosensor. The biosensor 10 includes a gate electrode 22, a source electrode 24, and a drain electrode 26 connected by a semiconductor material 28, and a microfluidic channel structure formed from a dielectric material 20. The microfluidic channel structure includes a first dielectric layer 32 between the gate 22 and the microfluidic channel 30 (where a first surface 34 is in contact with the microfluidic channel 30), a second dielectric layer 36 between the semiconductor material 28 and the microfluidic channel 30 (where a second surface 38 is in contact with the microfluidic channel 30), and a microfluidic channel 30 defined by one or more sidewalls 40 separating the first surface 34 from the second surface 38. As those skilled in the art will understand, the one or more sidewalls 40 may form a continuous sidewall defining the periphery of the microfluidic channel 30. The microfluidic channel structure also generally includes two or more openings indicated by 42, through which electrolyte fluid can be injected into or moved out of the microfluidic channel 30, or through which the electrolyte fluid can flow through the microfluidic channel 30, providing inlets and outlets.

[0068] In at least one embodiment, the gate electrode 22 is located at the bottom of the biosensor, as shown in Figures 1 and 2, so that the biosensor would be positioned during use. Figure 3A schematically illustrates an alternative embodiment that includes a gate substrate 44 supporting the gate electrode 22. Another alternative embodiment schematically illustrated in Figure 3B further includes a source electrode 24, a drain electrode 26, and a source / drain substrate 46 supporting the semiconductor material 28. In further alternative embodiments schematically illustrated in Figures 4A and 4B, the gate electrode is located at the top of the biosensor, so that the biosensor would be positioned during use. Figure 5 schematically illustrates an embodiment in which the gate electrode 22, microfluidic channel 30, source electrode 24, drain electrode 26, and semiconductor material 28 are supported by a substrate 48 prepared from a dielectric material 20 and oriented perpendicular to the directions shown in Figures 1-4B.

[0069] In such an embodiment, the distance between the gate electrode 22 and the semiconductor material 28 through the microfluidic channel 30 can be relatively small compared to the area of ​​the gate electrode, the semiconductor material, and the microfluidic channel. Without being bound by theory, such an embodiment is intended to provide a small capacitive distance between the gate electrode and the semiconductor material, and a relatively large conductive channel between the source electrode and the drain electrode, while providing a relatively small diffusion distance between the first surface and the second surface. In at least one embodiment, the area of ​​the biosensor device is about 500 μm². 2 From approximately 1000 μm 2 The range can extend to approximately 20 μm. In at least one embodiment, the length of the conductive channel formed between the source electrode and the drain electrode through the semiconductor material can range from approximately 20 μm to approximately 200 μm.

[0070] In at least one embodiment, the thickness of the first dielectric layer between the gate electrode and the first surface may range from about 50 nm to 500 nm. In at least one embodiment, the thickness of the first dielectric layer between the gate electrode and the first surface is about 200 nm. In at least one embodiment, the thickness of the second dielectric layer between the semiconductor and the second surface may range from about 50 nm to 500 nm. In at least one embodiment, the thickness of the second dielectric layer between the semiconductor and the second surface is about 200 nm.

[0071] In at least one embodiment, the distance between the first surface 34 and the second surface 38 in the microfluidic channel 30 can range from about 100 μm to about 1000 μm. In at least one embodiment, the distance between the first surface 34 and the second surface 38 in the microfluidic channel 30 can range from about 100 μm to about 600 μm. In at least one embodiment, the distance between the first surface 34 and the second surface 38 in the microfluidic channel 30 can range from about 300 μm to about 600 μm. In at least one embodiment, the distance between the first surface 34 and the second surface 38 in the microfluidic channel 30 can range from about 400 μm to about 600 μm. In at least one embodiment, the distance between the first surface 34 and the second surface 38 in the microfluidic channel 30 is about 300 μm. In at least one embodiment, the distance between the first surface 34 and the second surface 38 in the microfluidic channel 30 is about 400 μm. In at least one embodiment, the distance between the first surface 34 and the second surface 38 in the microfluidic channel 30 is about 500 μm. In at least one embodiment, the distance between the first surface 34 and the second surface 38 in the microfluidic channel 30 is about 550 μm. In at least one embodiment, the distance between the first surface 34 and the second surface 38 in the microfluidic channel 30 is about 600 μm.

[0072] Figures 6A-6C schematically illustrate an embodiment of the biosensor in which several source electrodes 24 and drain electrodes 26 are attached to a substrate 48 prepared from a dielectric material 20. The microfluidic channel structure is formed from the dielectric material 20 and includes a meandering microfluidic channel 30, a second dielectric layer 36 separating the semiconductor material 28 from the microfluidic channel 30, and a first dielectric layer 32 separating the microfluidic channel 30 from the gate electrode 22 attached to the outer surface of the first dielectric layer 32. It will be apparent to those skilled in the art that the individual source electrodes can be electrically connected to each other so that they act together as a single source electrode. Similarly, the individual drain electrodes can be electrically connected to each other so that they act together as a single drain electrode.

[0073] In at least one embodiment of the biosensor, the microfluidic channel may be positioned adjacent to the semiconductor material or adjacent to the gate electrode. In such embodiment, the semiconductor material and the gate electrode are separated from each other by a single dielectric layer, and changes in the capacitance of the microfluidic channel may result in edge or peripheral capacitance effects in the charge distribution between the gate electrode and the semiconductor material 28, which may be reflected in the output current of the field-effect transistor, as will be understood by those skilled in the art.

[0074] In at least one embodiment of the biosensor, the electrolyte gating function can be physically separated from the transistor function. Thus, in at least one such embodiment, a first dielectric layer of the microfluidic channel structure separates the first plate of the capacitor from the microfluidic channel, and a second dielectric layer separates the second plate of the capacitor from the microfluidic channel, thereby separating the first and second plates of the capacitor from each other by the first dielectric layer, the microfluidic channel, and the second dielectric layer. In such embodiment, as is well known in the art, either the first or second plate of the capacitor can be electrically connected to the gate electrode of a field-effect transistor, which includes a gate electrode, a source electrode, a drain electrode, a semiconductor material connecting the source and drain electrodes, and a dielectric layer between the gate electrode and the semiconductor material. In this way, changes in the capacitor voltage related to changes in the capacitance of the microfluidic channel can be reflected in the voltage of the gate electrode of the field-effect transistor and, consequently, its output current.

[0075] Such embodiments are schematically illustrated in Figures 7A and 7B. The illustrated embodiments include a microfluidic channel structure formed from a dielectric material 20, comprising a first dielectric layer 50, a second dielectric layer 52, and one or more sidewalls 53, which together define a microfluidic channel 30. The microfluidic channel structure is located between a first capacitor plate 54 and a second capacitor plate 56, thereby separating the microfluidic channel 30 from the first capacitor plate 54 by the first dielectric layer 50 and from the second capacitor plate 56 by the second dielectric layer 52. Either the first capacitor plate 54 (Figure 7A) or the second capacitor plate 56 (Figure 7B) may be electrically connected to the gate electrode of a field-effect transistor generally indicated by 58, so that changes in the capacitance and / or voltage of the capacitor can affect the gate voltage and, consequently, the output current of the field-effect transistor.

[0076] As is well understood in the art, the gate, source, and drain electrodes are made of any suitable conductive material. In at least one embodiment, the conductive material is a metal, including but not limited to aluminum, chromium, gold, platinum, and other metals known in the art. In at least one embodiment, the conductive material is a nonmetallic conductive element, including but not limited to graphite, graphene, carbon nanotubes, and other conductive allotropes of carbon. In at least one embodiment, the conductive material is a conductive compound, including but not limited to a conductive polymer. Suitable conductive polymers are well known in the art and include, but are not limited to, polyacetylene, polyphenylene vinylene, polypyrrole, polythiophene, polyaniline, polyphenylene sulfide, its copolymers, and other suitable polymers well known to those skilled in the art. In at least one embodiment, the conductive polymer is poly(3,4-ethylenedioxythiophene) polystyrene sulfonate (PEDOT:PSS).

[0077] In at least one embodiment, the gate electrode, source electrode, and drain electrode are attached to the substrate by any preferred technique known in the art, including but not limited to spin coating, deposition from solution, and printing, including but not limited to screen printing, inkjet printing, aerosol jet printing, transfer foil printing, and shadow mask printing. In at least one embodiment, the substrate is a flexible imide film, such as Kapton® film. In at least one embodiment, the substrate is the outer upper surface of either the first or second dielectric layer of the microfluidic channel structure. For example, in the embodiment shown in Figure 3A, the source electrode 24, drain electrode 26, and semiconductor material 28 may be attached to the outer surface of the second dielectric layer 36. Similarly, in the arrangement shown in Figure 4A, the gate electrode 22 may be attached to the outer surface of the first dielectric layer 32.

[0078] In at least one embodiment, the thickness of the gate electrode, source electrode, and drain electrode attached to the substrate may range from about 100 nm to about 10 μm. In at least one embodiment, the thickness of the gate electrode, source electrode, and drain electrode attached to the substrate may range from about 100 nm to about 5 μm. In at least one embodiment in which the gate electrode, source electrode, and drain electrode include metal, the thickness of the electrodes attached to the substrate may range from about 100 nm to about 500 nm. In at least one embodiment in which the gate electrode, source electrode, and drain electrode include a conductive polymer, the thickness of the electrodes attached to the substrate may range from about 200 nm to about 5 μm. A person skilled in the art can select electrode shapes, sizes, and relative positions suitable for a particular embodiment and can design and prepare such electrodes considering the teachings herein.

[0079] In at least one aspect, the semiconductor material includes an inorganic semiconductor. In at least one aspect, the inorganic semiconductor is silicon doped with an element such as boron, aluminum, gallium, indium, phosphorus, arsenic, or antimony. In at least one aspect, the inorganic semiconductor is indium gallium zinc oxide (IGZO), antimony tin oxide (ATO), nickel oxide (e.g., NiO x ), titanium oxide (e.g., TiO x ), indium oxide (e.g., InO x ), zinc oxide (e.g., ZnO x ), tin oxide (e.g., SnO x ), copper oxide (e.g., CuO x ), copper aluminum oxide (e.g., CuAlO x ), chromium oxide (e.g., CrO x ), cobalt oxide (e.g., CoO x ), tungsten oxide (e.g., WO x ), cerium oxide (e.g., CeO x ), niobium oxide (e.g., NbO x ), molybdenum oxide (e.g., MoO x ), silver oxide (e.g., AgO x ), gallium oxide (e.g., GaO x ), and other semiconductor metal oxides well known in the art, including but not limited to these semiconductor metal oxides.

[0080] In at least one embodiment, the semiconductor material includes any suitable organic or carbon-based semiconductor material known in the art, including but not limited to polycyclic aromatic hydrocarbons (such as rubrene, anthracene, and pentacene), metal phthalocyanines, carbon nanotubes, conjugated polymers (such as polythiophene, polyphenylene vinylene, and carbazole-dithiophene-benzothiadiazole copolymers), and block copolymers of such conjugated polymers. In at least one embodiment, the organic or carbon-based semiconductor includes, but is not limited to, 6,13-bis(triisopropylsilylethynyl)pentacene (TIPS-pentacene), polythiophenes and their copolymers (including, but not limited to, poly(3-hexylthiophene) (P3HT), 1,4-diketopyrrolo[3,4-c]pyrrole-thieno[3,2-b]thiophene (DPP-TTT), and dinaphthothienothiophene (DNTT)), metal phthalocyanines (including, but not limited to, copper phthalocyanine (CuPC)), and single-walled carbon nanotubes.

[0081] In at least one embodiment, the semiconductor material is deposited on a highly ordered material layer. Without being bound by theory, when the semiconductor material is an organic or carbon-based semiconductor material, the orientation of the formed organic semiconductor crystals or carbon nanotubes can be better controlled, resulting in better device-to-device consistency and improved charge mobility through the organic or carbon-based semiconductor material. In at least one such embodiment, the highly ordered material layer is a self-assembled monolayer. In at least one embodiment, the monolayer is a layer of octadecyltrichlorosilane with a thickness of one molecule.

[0082] The microfluidic channel structure, comprising first and second dielectric layers, comprises a dielectric material. In at least one embodiment, the dielectric material is a thermally meltable thermoplastic material that can be easily molded by 3D printing or extrusion and bonded to a separate structure prepared from the thermoplastic material. In at least one embodiment, the thermoplastic material surfaces to be bonded to each other may be bonded by means of partially melting the materials at the bonding surfaces, crosslinking, or using a solvent to partially dissolve the materials at the bonding surfaces. In at least one embodiment, the dielectric material may be annealed to reduce structural defects and improve dimensional stability, resulting in better consistency between devices during manufacturing. Suitable dielectric materials include, but are not limited to, thermoplastic polymers and their copolymers, including, but not limited to, polydienes, polycarbonates, polyimides, polyamides, polyacrylates, polyethers, polyurethanes, polyketones, polyhalodienes, polysiloxanes, polyolefins, fluoropolymers, polyesters, polysaccharides, and polyvinyls. In at least one embodiment, the dielectric material includes, but is not limited to, polylactic acid (PLA), polycaprolactone (PCL), poly(ethyl methacrylate) (PEMA), polybutylene terephthalate (PBT), polyethylene terephthalate (PET), polyethylene (PE), polypropylene (PP), polystyrene (PS), polytetrafluoroethylene (PTFE, Teflon®), cellulose and acylated cellulose derivatives (such as cellulose acetate, cellulose acetate butyrate, and cellulose propionate), acrylonitrile butadiene styrene (ABS), acrylonitrile styrene acrylate (ASA), polycaprolactam (PC), polybutadiene (PBD), polydimethylsiloxane (PDMS), nylon polyamide, and other polymers well known in the art.In at least one embodiment, the dielectric material includes, but is not limited to, polylactic acid (PLA), polycaprolactone (PCL), poly(ethyl methacrylate) (PEMA), polybutylene terephthalate (PBT), polyethylene terephthalate (PET), polyethylene (PE), polypropylene (PP), polystyrene (PS), polytetrafluoroethylene (PTFE), cellulose, cellulose acetate, cellulose butyrate acetate, cellulose propionate, acrylonitrile butadiene styrene (ABS), acrylonitrile styrene acrylate (ASA), polycaprolactam (PC), polybutadiene (PBD), and nylon polyamide.

[0083] In at least one embodiment, the dielectric material is polylactic acid (PLA). Without being bound by theory, PLA is intended to be biocompatible, hydrophobic, resistant to swelling upon contact with aqueous fluids, and functionalizable with antibodies, further details of which will be discussed later. In addition, PLA can be 3D printed and annealed, interacts well with metals and other polymer materials, and the complexity of the fabrication process can be reduced. Furthermore, PLA possesses dielectric and resistive properties that make it suitable for use in electrolyte-gate field-effect transistors.

[0084] In at least one embodiment, the dielectric material is polycaprolactone (PCL). Without being bound by theory, PCL is intended to be biodegradable, hydrophobic, resistant to swelling upon contact with aqueous fluids, easily processed at relatively low temperatures, possess dielectric and resistive properties suitable for use in electrolyte-gate field-effect transistors, and be readily covalently bonded to biorecognition bodies containing free amino groups, such as aptamers or antibodies. Further details will be described later.

[0085] In at least one embodiment, each of the first dielectric layer and the second dielectric layer comprises polylactic acid (PLA) and polycaprolactone (PCL), or a mixture of PLA and PCL. In at least one embodiment, each of the first dielectric layer and the second dielectric layer comprises PCL. In at least one embodiment, each of the first dielectric layer and the second dielectric layer comprises PLA. In at least one embodiment, the first dielectric layer comprises PCL and the second dielectric layer comprises PLA. In at least one embodiment, the first dielectric layer comprises PLA and the second dielectric layer comprises PCL.

[0086] In at least one embodiment, the first dielectric layer and / or the second dielectric layer may comprise multiple layers, each containing a different material. In at least one embodiment, the second dielectric layer may comprise a layer of PLA and a layer of polyvinyl alcohol (PVA) to which a monolayer of octadecyltrichlorosilane (OTS) is attached. As described above, the presence of the OTS monolayer may provide a base to which organic or carbon-based semiconductor materials can be attached, while improving the control of the orientation of organic semiconductor crystals or carbon nanotubes, resulting in improved consistency between devices.

[0087] In at least one embodiment, at least one of the first dielectric layer and the second dielectric layer may comprise a three-layer dielectric material comprising PLA, polyvinyl alcohol (PVA) containing cellulose nanocrystals, and thermally crosslinked PCL. In at least one embodiment, the second dielectric layer may comprise a three-layer dielectric material comprising PLA, polyvinyl alcohol (PVA) containing cellulose nanocrystals, and thermally crosslinked PCL. In at least one such embodiment, the three-layer material comprises a layer of PLA on which a semiconductor material can be deposited as described herein, a layer of PVA containing cellulose nanocrystals deposited on the opposite side of the PLA layer on which a semiconductor material can be deposited, and a layer of PCL on the opposite side of the PVA-cellulose layer bonded to the PLA layer, thereby the PCL layer forming a second surface of the second dielectric layer and being bonded to a biorecognition body as described herein. Without being bound by theory, embodiments of the device comprising such three-layer dielectric layers are intended to have a longer shelf life than embodiments lacking the three-layer dielectric layers.

[0088] In at least one embodiment, at least one of the first dielectric layer and the second dielectric layer includes a coating layer on at least one of the first and second surfaces. In at least one such embodiment, the coating layer is prepared from a material that facilitates bonding at least one of the first and second surfaces to a biorecognition body, as discussed below. Alternatively, at least one of the first and second dielectric layers may be prepared from such a material, facilitating bonding of the respective first or second surfaces to a biorecognition body. Such materials are well known in the art and include, but are not limited to, fluoropolymers such as polydimethylsiloxane (PDMS), poly(methyl methacrylate) (PMMA), and CYTOP®, Teflon®, and polyvinylidene fluoride (PVDF).

[0089] The microfluidic channel is configured to receive an electrolyte fluid such that the electrolyte fluid is in contact with an inner surface including first and second surfaces. In at least one embodiment, the electrolyte fluid is a sample obtained from an environmental source. In at least one embodiment, the electrolyte fluid is a sample obtained from an agricultural source. In at least one embodiment, the electrolyte fluid is a sample obtained from a food source. In at least one embodiment, the electrolyte fluid is a sample obtained from a manufacturing source. In at least one embodiment, the electrolyte fluid is a biological sample or fluid including, but not limited to, blood, plasma, urine, sweat, interstitial fluid, cerebrospinal fluid, exhaled condensate, and saliva. In at least one embodiment, the electrolyte fluid is saliva. In at least one embodiment, the electrolyte fluid is blood. In at least one embodiment, the electrolyte fluid is plasma. In at least one embodiment, the electrolyte fluid is a control fluid, which may also be used for calibrating a biosensor. The microfluidic channel may have any shape or configuration that causes the channel to receive the electrolyte fluid such that the electrolyte fluid is in contact with an inner surface. Preferred configurations for microfluidic channels include, but are not limited to, those shown in Figures 8A to 8D. Other preferred configurations will be readily identified and prepared by those skilled in the art.

[0090] In at least one embodiment, the biosensor includes a calibration well for use in calibrating and standardizing the output current of the biosensor. An embodiment of the microfluidic channel including the sample well and the calibration well is shown in Figure 8B. In at least one embodiment, the calibration well contains a reference electrolyte of a known concentration. In at least one embodiment, measurement of the output current when the reference electrolyte is present in the microfluidic channel enables calibration of the device and normalization of intra-device variations.

[0091] At least one of the first surface 34 and the second surface 38 within the microfluidic channel is responsible for the biorecognition body that recognizes and interacts with the analyte in the electrolyte fluid. Without being bound by theory, this interaction is thought to create a charge separation region in addition to the electrolytic bilayer at the solid dielectric-electrolyte fluid interface, resulting in a change in capacitance from the device gate to the channel, which in turn leads to a measurable change in the biosensor's output current that can reflect the presence or concentration of the analyte.

[0092] Suitable biorecognition entities are well known in the art and include, but are not limited to, antibodies, antigens, enzymes, receptors, aptamers, ligands, etc. In at least one embodiment, the biorecognition entity is an aptamer. As used herein and as known in the art, the term “aptamer” is intended to mean a single-stranded polydeoxyribonucleotide or polyribonucleotide that folds into a defined three-dimensional structure and is capable of recognizing and binding to a specific target, including, but not limited to, proteins or small molecules. In at least one embodiment, the aptamer may include naturally occurring nucleotide residues (i.e., as understood in the art, a ribose or deoxyribose moiety having one of the naturally occurring bases adenine, thymine, cytosine, guanine, or uracil, linked to an adjacent nucleotide residue via a monophosphate bond between the 5' hydroxyl group of one ribose or deoxyribose moiety and the 3' hydroxyl group of the adjacent ribose or deoxyribose moiety). In at least one embodiment, the aptamer may comprise a non-natural nucleotide residue or a chemically modified nucleotide residue, as understood in the art. In at least one embodiment, the biorecognition entity is an antibody.

[0093] In at least one embodiment, the analyte is a biomarker whose presence or concentration indicates a medical condition. Preferred analytes include, but are not limited to, drugs, metabolites, hormones, neurotransmitters, enzymes, carcinogens, peptides, proteins, electrolytes, metal ions, nucleic acids, cells, and other chemical or biochemical entities that are preferably measured in an electrolyte fluid, including but not limited to biological fluids. In at least one embodiment, the analyte is an analyte useful for measuring nerve function or detecting neurodegenerative diseases, including but not limited to dopamine, α-synuclein, β-amyloid, and tau protein. In at least one embodiment, the analytes useful for measuring nerve function or detecting neurodegenerative diseases are one or more of dopamine and α-synuclein, which are useful analytes for detecting Parkinson's disease. In at least one embodiment, the analytes useful for measuring nerve function or detecting neurodegenerative diseases are one or more of β-amyloid and tau protein, which are useful analytes for detecting Alzheimer's disease.

[0094] In at least one embodiment, the analyte is useful for measuring cardiac function or detecting cardiac disease or disorder, and includes, but is not limited to, B-type natriuretic peptide (BNP), N-terminal pro-B-type natriuretic peptide (NT-proBNP), troponin, and creatine kinase. In at least one embodiment, the analyte is useful for detecting blood disease or disorder, and includes, but is not limited to, bilirubin and ferritin. In at least one embodiment, the analyte is useful for detecting disease or disorder associated with infection, inflammation, or stress, and includes, but is not limited to, cortisol, lactate, C-reactive protein, interleukin, tumor necrosis factor (TNF), procalcitonin, and inflammatory cytokines. In at least one embodiment, the analyte is useful for measuring liver function or detecting liver disease or disorder, and includes, but is not limited to, bilirubin, alanine aminotransferase, aspartate aminotransferase, gamma-glutamyltransferase (GGT), albumin, and carcinoembryonic antigen. In at least one embodiment, the analyte is useful for measuring renal function or detecting kidney disease or impairment, and includes, but is not limited to, uric acid, creatinine, and urea. In at least one embodiment, the analyte is useful for detecting cancer or precancerous symptoms, and includes, but is not limited to, tumor necrosis factor (TNF), prostate-specific antigen (PSA), carcinoembryonic antigen, tumor-associated antigen, and inflammatory cytokines.

[0095] As used herein, and unless otherwise specified, the terms “bear” and “bearing” are intended to mean, when used in relation to a surface bearing a body, including but not limited to a biorecognition body, that the body is attached to or incorporated into the surface so that it is exposed to a liquid in contact with the surface, but is not readily removed from or washed away by contact with such fluid. In at least one embodiment, the biorecognition body is covalently or non-covalently bonded to the surface. In at least one embodiment, the biorecognition body is covalently or non-covalently bonded to at least one intermediate body, which itself is covalently or non-covalently bonded to the surface. In at least one embodiment, the surface integrally includes at least one of the at least one intermediate body, which is covalently or non-covalently bonded to a biorecognition body. As a non-limiting example, the surface may incorporate or bond to a biorecognition body capable of directly recognizing and bonding to an analyte, or a biorecognition body capable of recognizing and bonding to at least one additional biorecognition body, which itself directly recognizes and bonds to an analyte.

[0096] Figures 9A to 9C schematically illustrate the recognition and binding of the analyte 62 by the biorecognition body 60. For convenience, Figures 9A to 9C indicate that the biorecognition body 60 is located within the microfluidic channels 30 on the first surface 34 of the first dielectric layer 32. However, it is also intended that the biorecognition body 60 may be located within the microfluidic channels 30 on the second surface 38 of the second dielectric layer 36, or on both the first surface 34 and the second surface 38.

[0097] Figure 9A schematically illustrates how the free analyte 62 is directly recognized and bound to the biorecognition body 60, forming the bound analyte 64. Figure 9B schematically illustrates a competitive binding mode in which the biorecognition body 60 first binds to a competing ligand 66, which is then replaced by the analyte 62 to form the bound analyte 64. Figure 9C schematically illustrates a binding mode in which both the biorecognition body 60 and the analyte 62 recognize and bind to the secondary body 68, forming the bound analyte 64. As will be apparent to those skilled in the art, the secondary body 68 may also be a ligand that is recognized and binds to the biorecognition body 60 and directly binds to the analyte 62 itself, that is, it facilitates the binding of the biorecognition body 60 to the analyte 62. Alternatively, the secondary body 68 may be a biorecognition body that recognizes and binds to the analyte 62 itself and also recognizes and binds to the biorecognition body 60, or is recognized and bound by the biorecognition body 60. Other coupling modes will be obvious to those skilled in the art from the viewpoint of the teachings described herein and will be easily carried out.

[0098] At least one of the first surface and the second surface may be configured to carry a biorecognition body by methods well known in the art. In at least one embodiment, when the biorecognition body is an aptamer and at least one of the first surface and the second surface contains poly(methyl methacrylate) (PMMA), the aptamer can be immobilized on the PMMA surface by UV activation, as described in "RS Massey and R. Prakash, "A Low-Temperature-Processed, Soft-Fluidic OEGFET Saliva Aptasensor for Cortisol," in IEEE Journal on Flexible Electronics, vol. 1, no. 1, pp. 64-72, Jan. 2022". In at least one embodiment, when at least one of the first and second surfaces contains polycaprolactone (PCL), the biorecognition body, which includes but is not limited to aptamers or antibodies, can be bonded to the first or second surface using the reagents 1-ethyl-3-(3-dimethylaminopropyl)carbodiimide (EDC) and N-hydroxysuccinimide (NHS) using chemical procedures well known in the art and, for example, as described in "M. Cooper, “Sensor surfaces and receptor deposition,” in Label-free biosensors techniques and applications, M. Cooper Ed. Cambridge, UK: Cambridge University Press, pp. 110 - 142, 2009," thereby covalently bonding amine groups on the biorecognition body to polycaprolactone. In at least one embodiment, when at least one of the first and second surfaces contains polylactic acid (PLA), the biorecognition body contains a polylactic acid-polyethylene glycol-carboxylic acid block copolymer, regardless of whether the first or second surface is the biorecognition body.The presence of such block copolymers containing carboxyl groups can facilitate covalent bonding to a first or second surface of a biorecognition body, including but not limited to aptamers or antibodies, by crosslinking using the EDC-NHS method described above.

[0099] Another aspect of this application provides a method for producing an organic electrolyte gate field-effect transistor biosensor as described herein. In at least one embodiment, the method is as follows: 3D printing a thermoplastic dielectric material to form a microfluidic channel structure containing at least one inner surface; A microfluidic channel structure is placed between the gate electrode and the semiconductor material that electrically contacts the source electrode and drain electrode, such that at least one inner surface is isolated from the gate electrode and semiconductor material by a dielectric material; and, The biorecognition unit is coupled to at least a part of at least one inner surface. Includes.

[0100] In at least one embodiment, a thermoplastic dielectric material is 3D printed to form a first dielectric layer as described herein between a gate electrode and at least one internal surface. In at least one embodiment, a thermoplastic dielectric material is 3D printed to form a second dielectric layer as described herein between a semiconductor material and at least one internal surface.

[0101] In at least one embodiment, the method is as follows: Attaching the gate electrode to the first substrate; 3D printing a dielectric material onto a gate electrode to form a first dielectric layer having a first surface; The source electrode and drain electrode are attached to the second substrate; Deposition of a semiconductor material onto a second substrate so as to electrically connect the source electrode and drain electrode to the semiconductor material; 3D printing a dielectric material onto a semiconductor material to form a second dielectric layer having a second surface; The biorecognition body is attached to at least one of the first surface and the second surface; and A first dielectric layer is bonded to a second dielectric layer having one or more sidewalls, thereby defining a microfluidic channel between the first surface, the second surface, and one or more sidewalls. Includes.

[0102] In at least one embodiment, the method is as follows: Attaching source and drain electrodes to a substrate; Deposition of a semiconductor material onto a substrate so that the source electrode and drain electrode are connected to the semiconductor material; A dielectric material is 3D printed onto a semiconductor material to form a microfluidic channel structure, the microfluidic channel structure comprising a first dielectric layer having a first surface, a second dielectric layer having a second surface, and one or more sidewalls between the first and second dielectric layers to define a microfluidic channel between the first and second surfaces, wherein the second dielectric layer is between the semiconductor material and the microfluidic channel; The biorecognition body is attached to at least one of the first surface and the second surface; and The gate electrode is deposited on the first dielectric layer such that the first dielectric layer is between the gate electrode and the microfluidic channel. Includes.

[0103] In at least one embodiment, the method is as follows: Attaching a gate electrode to a substrate; A dielectric material is 3D printed onto the gate electrode to form a microfluidic channel structure, the microfluidic channel structure comprising a first dielectric layer having a first surface, a second dielectric layer having a second surface, and one or more sidewalls between the first and second dielectric layers to define a microfluidic channel between the first and second surfaces, wherein the first dielectric layer is located between the gate electrode and the microfluidic channel; The biorecognition body is attached to at least one of the first surface and the second surface; A semiconductor material is deposited on a second dielectric layer, where the second dielectric layer is located between the semiconductor material and the microfluidic channel; The source electrode and drain electrode are attached to a second dielectric layer so as to electrically connect them to the semiconductor material. Includes.

[0104] In at least one embodiment, the method includes depositing a self-assembled monolayer onto a second substrate or a second dielectric layer, and depositing an organic or carbon-based semiconductor material onto the self-assembled monolayer. In at least one embodiment, the self-assembled monolayer comprises octadecyltrichlorosilane.

[0105] In at least one embodiment, the microfluidic channel structure is three-dimensionally printed from a dielectric material to form a monolithic structure. By fabricating the biosensor in this way, the first and second dielectric layers defining the microfluidic channels can be prepared with greater reproducibility during the 3D printing process and formed uniformly within the microfluidic channel structure, avoiding the problems associated with molding individual microfluidic channels from PDMS and bridging them to individually formed first and second dielectric layers, as seen in previously known devices.

[0106] In another aspect, this application provides a method for analyzing an electrolyte fluid for an analyte, comprising exposing an organic electrolyte gate field-effect transistor biosensor as described herein to the electrolyte fluid and determining one or more characteristics of the analyte. In at least one embodiment, the electrolyte fluid is a sample obtained from an environmental source. In at least one embodiment, the electrolyte fluid is a sample obtained from an agricultural source. In at least one embodiment, the electrolyte fluid is a sample obtained from a food source. In at least one embodiment, the electrolyte fluid is a sample obtained from a manufacturing source. In at least one embodiment, the electrolyte fluid is a biological fluid. In at least one embodiment, the biological fluid is one or more of blood, plasma, urine, sweat, interstitial fluid, cerebrospinal fluid, exhaled condensate, or saliva. In at least one embodiment, the electrolyte fluid is saliva. In at least one embodiment, the electrolyte fluid is blood. In at least one embodiment, the electrolyte fluid is plasma. In at least one embodiment, the analyte is a biomarker whose presence or concentration indicates a medical condition, as described herein. In at least one embodiment, exposure of the biosensor to an electrolyte fluid includes exposure of the microfluidic channels of the biosensor to the electrolyte fluid. In at least one embodiment, determining one or more features of the analyte includes determining the presence of the analyte in the electrolyte fluid or measuring the concentration of the analyte in the electrolyte fluid. In at least one embodiment, the interaction of the analyte with the biorecognition body results in the formation of a charge separation region at the interface between the electrolyte fluid and at least one of the first dielectric layer and the second dielectric layer, a change in the net capacitance of the gate-to-channel of the biosensor, and a change in the output current of the biosensor, where the change in output current indicates one or more features of the analyte in the electrolyte fluid.

[0107] When used herein, and unless otherwise specified, the terms “a” and “an” are intended to include both plural and singular forms and may be interpreted as meaning “one or more.”

[0108] When used herein, the terms “about” or “approximately” applied to numerical values ​​or ranges of values ​​are intended to mean that, given the nature or precision of the measurement, the stated value may vary within an acceptable margin of error with respect to the measured quantity, and that such variation is considered equivalent to the stated value in the art and provides the same function or result. For example, the degree of error may be indicated by the number of significant numbers provided for the measurement, as understood in the art, and may include, but is not limited to, a variation of ±1 in the most accurate significant number reported for the measurement. A typical exemplary degree of error is within 20%, preferably within 10%, and more preferably within 5%, of a given value or range of values. Alternatively, particularly in biological systems, the terms “about” and “approximately” may mean a value within one order of magnitude, preferably within five times, and more preferably within two times, of a given value. Numerical values ​​described herein are approximate unless otherwise stated, and the terms “about” or “approximately” may be implied, even when not explicitly stated.

[0109] As used herein, the term “substantially” refers to the complete or near-complete range or degree of an action, feature, characteristic, state, structure, item, or result. For example, an object that is “substantially” in a given position, including but not limited to being vertical, horizontal, or adjacent to or aligned with another object, means that the object is either completely or nearly completely in that position. The degree of strictly permissible deviation from absolute completeness may, in some cases, depend on the particular context. However, generally speaking, approaching completion would be because it would have the same overall result as if absolute and complete completion had been achieved.

[0110] The use of “substantially” applies equally when used in a negative sense, referring to the complete or near-complete absence of an action, feature, characteristic, state, structure, item, or result. For example, a composition that “substantially” does not contain an ingredient or element will either completely lack that ingredient or element, or nearly completely lack it, and will therefore have the same effect as if it were completely lacking. In other words, a composition that “substantially” does not contain an ingredient or element may still actually contain such item, unless its effect is measurable or significant.

[0111] When used herein, terms indicating relative directions or orientations, including but not limited to "top," "bottom," "upper surface," "lower surface," "vertical," "horizontal," "outside," "inside," "front," and "back," are intended to facilitate the description of the invention by indicating relative directions or orientations in ordinary use, and are not intended to limit the scope of the invention to any such directions or orientations.

[0112] example Other features of the present invention will become apparent from the following non-limiting examples illustrating the principle of the present invention.

[0113] Example 1: General procedure for preparing a biosensor Specific examples of this biosensor can be prepared using the following general procedure: The gate electrode is printed onto a flexible Kapton® substrate with silver nanoparticle ink using screen printing or extrusion printing techniques. A solution of 5% by weight polycaprolactone (PCL) dissolved in chloroform is deposited onto the gate electrode by extrusion printing and dried at 30°C. The microfluidic channel structure can be formed by 3D printing polylactic acid (PLA) onto the PCL surface or PCL surface. In a specific example containing PLA, the PCL surface was activated with 0.1 M NaOH, and PLA was deposited onto a portion of the PCL surface by 3D printing from a molten state (200°C) to form microfluidic channels, leaving a portion of the PCL surface exposed.

[0114] Next, the biorecognition body is covalently bound to the exposed PCL surface using 1-ethyl-3-(3-dimethylaminopropyl)carbodiimide (EDC) and N-hydroxysuccinimide (NHS) reagents, for example, as described in "M. Cooper, “Sensor surfaces and receptor deposition,” in Label-free biosensors techniques and applications, M. Cooper Ed. Cambridge, UK: Cambridge University Press, pp. 110 - 142, 2009". Alternatively, a poly(methyl methacrylate) (PMMA) coating may be applied to the PCL surface, and then the biorecognition body may be covalently bound to the PMMA coating. In at least one embodiment, the aptamer may be bound to the PMMA coating by UV-activated drop casting, as described in "RS Massey and R. Prakash, IEEE Journal on Flexible Electronics (2022), 1(1): 64-72". The antibody may also be covalently bound to PMMA using the EDC-NHS procedure described above.

[0115] The source and drain electrodes are printed onto a flexible Kapton® substrate using silver nanoparticle ink, again using screen printing or extrusion printing techniques. The semiconductor material is deposited onto the flexible substrate as described in Example 2 below and annealed to connect the source and drain electrodes. A 9 wt% PLA solution dissolved in 1,4-dioxane, or a 5 wt% PCL solution dissolved in chloroform, is deposited onto the semiconductor material by extrusion printing and, while still wet, is brought into contact with the exposed surface of the 3D printed PLA or PCL microchannels to seal the microfluidic channels.

[0116] Alternatively, a triple dielectric layer can be formed on a semiconductor material. The PLA layer on the semiconductor material is deposited by spin-coating with a 2 mg / mL PLA solution dissolved in chloroform, followed immediately by spin-coating the PLA layer with an 80 mg / mL polyvinyl alcohol (PVA) solution containing 0.75 wt% cellulose nanocrystals. The PLA-PVA-cellulose bilayer is annealed at 150°C under vacuum for 1 hour, a 2 mg / mL solution of toluene diisocyanate-terminated polycaprolactone (TPCL) is spin-coated onto the PVA-cellulose layer, and this triple layer is annealed at 200°C under vacuum for 15 minutes.

[0117] A specific example of this biosensor, prepared according to the general procedure outlined above, was tested for transistor function using deionized water as the electrolyte in the microfluidic channel, employing a general procedure well-known in the art. Various voltages were applied to the source electrode, drain electrode, and gate electrode, and the device output current and device transfer function response (output current modulated by the gate bias voltage) were measured using standard laboratory equipment such as a semiconductor parameter analyzer, or a custom-made system-onboard hybrid integrated system for low-power portable analysis. Representative graphs obtained for this specific example of the biosensor are shown in Figures 10A to 10F, respectively.

[0118] Example 2: Deposition of semiconductor materials Various specific examples of this biosensor were prepared using alternative methods for depositing semiconductor material onto source and drain electrodes, in addition to using the general procedure of Example 1. Specific examples involve depositing the semiconductor material TIPS-pentacene by drop casting, spin welding, screen printing, or inkjet printing (piezoelectric drop-on-demand (DoD)) by a 1% by weight solution of TIPS-pentacene dissolved in chlorobenzene, or by inkjet printing a 1% by weight solution of TIPS-pentacene dissolved in 15% chloroform and 84% chlorobenzene.

[0119] Alternatively, semiconductor materials containing carbon nanotubes can be deposited onto a flexible substrate. A mixture of commercially available carbon nanotubes dispersed in toluene and poly(9,9'-didodecylfluorene-co-N-(2'-decyltetradecane)-carbazole (PCPF) (Mn=65, PD=2.7) is purified as described in "Rice, N. et al, Advanced Electronic Materials (2019), 5(1): 1800539" and deposited using drop casting or inkjet / aerosol jet printing.

[0120] Figure 11 shows the device output curves of alternative examples prepared by inkjet printing from a 1 wt% TIPS-pentacene solution dissolved in 15% chloroform and 84% chlorobenzene, or by screen printing, spin welding, and drop casting of semiconductor materials from a 1 wt% TIPS-pentacene solution dissolved in chlorobenzene. As can be seen from the data presented in Figure 11, inkjet printing of the semiconductor material from a 1 wt% TIPS-pentacene solution dissolved in 15% chloroform and 84% chlorobenzene provided devices with the most preferred characteristics. This method can provide precise adhesion and controlled device structure, which will improve device yield and reduce batch-to-batch variability.

[0121] Example 3: Detection test of analyte A specific example of this biosensor device was prepared according to the procedures of Examples 1 and 2, containing both PCL and PLA as dielectric materials and carrying an aptamer that recognizes either cortisol or α-synuclein. Figures 12-14 show the characteristics of the device measured in the presence of buffers containing various concentrations of cortisol or α-synuclein, or in the presence of saliva supplemented with various concentrations of α-synuclein, as described in "Massey, RS et al, “Non-invasive Monitoring of Alpha-synuclein in Saliva for Parkinson's Disease using Organic Electrolyte Gated FET Aptasensor”, ACS Sensors (2023), 8(8): 3116-3126". The results shown in Figures 12-14 demonstrate that the device exhibits characteristics that enable it to function as a biosensor.

[0122] Example 4: Biosensor containing carbon nanotube semiconductor material Specific examples of this biosensor device, in which the semiconductor material comprises single-walled carbon nanotubes, were prepared as described in Examples 1 and 2. In at least one such embodiment, the dielectric material deposited as a second dielectric layer between the semiconductor material and the PCL microfluidic channel structure is a three-layer dielectric material comprising PLA, PVA containing cellulose nanocrystals, and thermally crosslinked PCL, as described herein. Figures 15A–15D show the characteristics of such specific devices as measured when the microfluidic channel contains deionized water (Figures 15A and 15B) or a commercially available tris(tris(hydroxymethyl)aminomethane) acetate EDTA (ethylenediaminetetraacetic acid) (TAE) buffer (Figures 15C and 15D). As can be seen from Figures 15A–15C, the transconductance and channel current of the device at a given applied gate voltage increase as the conductivity of the electrolyte increases from that of deionized water to that of the buffer solution.

[0123] Example 5: A biosensor that uses antibodies as the biological recognition mechanism. As described in Examples 1 and 2, specific biosensor devices, each carrying either an aptamer that recognizes α-synuclein or an antibody that recognizes amyloid-β-42 as the biorecognition body, were prepared and tested for the detection of the corresponding analytes (α-synuclein or amyloid-β-42, respectively) as described in "Massey, RS et al, ACS Sensors (2023), 8(8): 3116-3126," except that the serum sample was diluted 1:1 with phosphate-buffered saline (PBS) instead of deionized water. As can be seen from the data shown in Figures 16A-F, when each analyte was present in TAE buffer (Figures 16A and 16B), in a real salivary supernatant sample (Figures 16C and 16D), and in a real serum sample (Figures 16E and 16F), the devices biofunctionalized with the antibody or aptamer exhibited equivalent concentration-dependent detection behavior. Aptamer-based sensors showed some differences from antibody-based sensors in terms of detection limit (LOD) and matrix effects. Without being bound by theory, these differences are intended to be at least in part due to the larger size of antibodies compared to typical aptamers, resulting in differences in Debye length on biofunctionalized surfaces and differences in the speed at which binding to the analyte saturates within the sensor region.

[0124] Example 6: Comparison with known devices Specific examples A to D of this biosensor were prepared according to Examples 1 and 2 described above. Specific examples A and C were prepared using only PCL to prepare the microfluidic channel structure, while the microfluidic channel structure of Specific Example B contains both PCL and PLA, and the microfluidic channel structure of Specific Example D contains PCL in the first dielectric layer between the gate and the microfluidic channel, and a three-layer dielectric layer containing PLA, PVA containing cellulose nanocrystals, and thermally crosslinked PCL in the second dielectric layer between the semiconductor material and the microfluidic channel, as described herein. Specific examples A and B contain TIPS-pentacene as the semiconductor material, while specific examples C and D contain carbon nanotubes as the semiconductor material. Comparative devices 1 and 2 were prepared according to the instructions in "(RS Massey and R. Prakash, "A Low-Temperature-Processed, Soft-Fluidic OEGFET Saliva Aptasensor for Cortisol," in IEEE Journal on Flexible Electronics, vol. 1, no. 1, pp. 64-72, Jan. 2022)" and "Massey, RS; Prakash, R. Modeling the Double Layer Capacitance Effect in Electrolyte Gated FETs with Gel and Aqueous Electrolytes. Micromachines (2021), 12, 1569, 1-15".

[0125] The characteristics of comparative devices 1 and 2, as well as the characteristics of specific examples A to D, were measured using standard procedures known in the relevant art, and the results are shown in Table 1 below. "W / L" represents the width-to-length ratio of the device, i.e., the aspect ratio. [Table 1]

[0126] As can be seen from the results presented in Table 1, Specific Examples A and B exhibit lower gate leakage current and a higher on / off ratio compared to comparative devices 1 and 2, which is expected to result in improvements in the sensitivity and shelf life of this biosensor. In addition, Specific Examples C and D show performance at least comparable to comparative devices 1 and 2, but demonstrate better protection of the semiconductor layer from moisture and environmental doping, and show less batch-to-batch variability and sensor drift.

[0127] The embodiments described herein are intended to illustrate the composition and method and are not intended to limit the scope of the invention. Various modifications and changes consistent with the description as a whole and readily understandable to those skilled in the art are intended to be included. The appended claims should not be limited by any particular embodiment illustrated and should be interpreted in the broadest way consistent with the description as a whole.

Claims

1. An organic electrolyte gate-type field-effect transistor biosensor, wherein: Guard gates; Source electrode; Drain electrode; Semiconductor material in electrical contact with the source electrode and drain electrode; and, The microfluidic channel structure includes at least one thermoplastic dielectric material and at least one inner surface comprising a biorecognition body, wherein at least one inner surface is separated from the gate electrode, source electrode, drain electrode, and semiconductor material by the thermoplastic dielectric material. The aforementioned biosensor.

2. An organic electrolyte gate type field-effect transistor biosensor according to claim 1, wherein the thermoplastic dielectric material is one or more materials selected from the group consisting of polylactic acid (PLA), polycaprolactone (PCL), poly(ethyl methacrylate) (PEMA), polybutylene terephthalate (PBT), polyethylene terephthalate (PET), polyethylene (PE), polypropylene (PP), polystyrene (PS), polytetrafluoroethylene (PTFE), cellulose, acylated cellulose derivatives, cellulose acetate, cellulose acetate butyrate, cellulose propionate, acrylonitrile butadiene styrene (ABS), acrylonitrile styrene acrylate (ASA), polycaprolactam (PC), polybutadiene (PBD), polydimethylsiloxane (PDMS), and nylon polyamide. The aforementioned biosensor.

3. The organic electrolyte gate type field-effect transistor biosensor according to claim 2, wherein the thermoplastic dielectric material comprises at least one of polylactic acid (PLA) and polycaprolactone (PCL).

4. An organic electrolyte gate type field-effect transistor biosensor according to any one of claims 1 to 3, wherein the microfluidic channel structure is as follows: A first dielectric layer having a first surface; A second dielectric layer having a second surface parallel to and facing the first surface; and, To separate the first dielectric layer from the second dielectric layer, one or more side walls adjacent to the first dielectric layer and the second dielectric layer, Includes, Here, the first surface, the second surface, and one or more sidewalls define microfluidic channels between them, and the microfluidic channels are configured to receive electrolyte fluid such that the electrolyte fluid comes into contact with the first surface and the second surface, and further, Here, at least one of the first surface and the second surface is responsible for the biorecognition body. The aforementioned biosensor.

5. The organic electrolyte gate type field-effect transistor biosensor according to claim 4, wherein one or both of the first dielectric layer and the second dielectric layer comprises at least one of polylactic acid (PLA) and polycaprolactone (PCL).

6. The organic electrolyte gate type field-effect transistor biosensor according to claim 4 or 5, wherein at least one of the first dielectric layer and the second dielectric layer comprises a layer containing PLA, a layer containing polyvinyl alcohol (PVA) and nanocrystalline cellulose, and a layer containing PCL.

7. The organic electrolyte gate type field-effect transistor biosensor according to any one of claims 1 to 6, wherein the semiconductor material is an organic or carbon-based semiconductor material.

8. The organic electrolyte gate type field-effect transistor biosensor according to claim 7, wherein the organic or carbon-based semiconductor material is selected from the group consisting of 6,13-bis(triisopropylsilylethynyl)pentacene (TIPS-pentacene), polythiophenes and their copolymers, metal phthalocyanines, and single-walled carbon nanotubes.

9. The organic electrolyte gated field-effect transistor biosensor according to claim 8, wherein the organic or carbon-based semiconductor material is TIPS-pentacene or single-walled carbon nanotubes.

10. The organic electrolyte gate type field-effect transistor biosensor according to any one of claims 7 to 9, further comprising a self-assembled monolayer, wherein an organic or carbon-based semiconductor material is attached to the self-assembled monolayer.

11. The organic electrolyte gate type field-effect transistor biosensor according to any one of claims 1 to 10, wherein the biorecognition body is an aptamer or an antibody.

12. A method for analyzing an electrolyte fluid for an analyte, comprising exposing an organic electrolyte gate-type field-effect transistor biosensor according to any one of claims 1 to 11 to the electrolyte fluid and determining one or more characteristics of the analyte in the electrolyte fluid. The aforementioned method.

13. The method according to claim 12, wherein the electrolyte fluid is a biological fluid selected from blood, plasma, urine, sweat, interstitial fluid, cerebrospinal fluid, saliva, and exhaled condensate.

14. The method according to claim 13, wherein the electrolyte fluid is selected from blood, plasma, and saliva.

15. The method according to any one of claims 12 to 14, wherein one or more features of the analyte in the electrolyte fluid include the presence of the analyte in the electrolyte fluid or the concentration of the analyte in the electrolyte fluid.

16. The method according to any one of claims 12 to 15, wherein the analyte in the electrolyte fluid is selected from the group consisting of drugs, metabolites, hormones, neurotransmitters, enzymes, carcinogens, peptides, proteins, electrolytes, metal ions, nucleic acids, and cells.

17. The method according to any one of claims 12 to 16, wherein the analyte in the electrolyte fluid is an analyte useful for measuring nerve function, cardiac function, liver function or kidney function, or for detecting neurodegenerative diseases, heart disease or cardiac disorders, blood diseases or blood disorders, infectious diseases, inflammation or stress-related diseases, liver diseases or liver disorders, kidney diseases or kidney disorders, and cancer or precancerous symptoms.

18. The method according to claim 17, wherein the analyte in the electrolyte fluid is selected from the group consisting of α-synuclein, β-amyloid, and tau protein.

19. The method according to any one of claims 12 to 18, wherein the interaction of the analyte with the biorecognition body results in one or more of the following: the formation of a charge separation region at the interface between the electrolyte fluid and at least one of the first dielectric layer and the second dielectric layer; a change in the net capacitance of the gate-to-channel of the biosensor; and a change in the output current of the biosensor, wherein the change in the output current indicates one or more characteristics of the analyte in the electrolyte fluid.

20. A method for producing an organic electrolyte gate type field-effect transistor biosensor according to any one of claims 1 to 11, the following: 3D printing a thermoplastic dielectric material to form a microfluidic channel structure containing at least one inner surface; A microfluidic channel structure is placed between the gate electrode and the semiconductor material that electrically contacts the source electrode and drain electrode, such that at least one inner surface is isolated from the gate electrode and semiconductor material by a dielectric material; and, The biorecognition unit is coupled to at least a part of at least one inner surface. The method, including the method described above.

21. The method according to claim 20, wherein the microfluidic channel structure is printed three-dimensionally as a monolithic structure from a thermoplastic dielectric material.

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